Dual energy imaging with kinestatic charge detector

ABSTRACT

A method and apparatus for providing dual energy radiation images of a patient in a kinestatic charge detection system utilizes first and second kinestatic charge detectors commonly connected to apparatus for rotating the detectors about a radiation source at a predetermined velocity. The detectors are adjusted such that ion drift velocity in each detector is equal in magnitude but opposite in direction to the velocity of rotation of the detectors. The radiation from the radiation source is modulated such that relatively low average energy radiation impinging on the patient is received by the first detector and relatively high average energy radiation impinging on the patient is received by the second detector. As the radiation is scanned across the patient and received by the detectors, output data from each detector provides information representative of the intensity of radiation received during the scan. The output data is then combined by a imaging system to form a difference image.

This invention relates to ionization chamber x-ray detectors and, morespecifically, to a method and apparatus for obtaining a dual energydifference image in a kinestatic charge detector system.

BACKGROUND OF THE INVENTION

The optimal detection of ionizing radiation in two dimensions is thecentral problem in computed tomography, digital radiography, nuclearmedicine imaging and related disciplines. Many different types ofdetectors (e.g., non-electronic, analog electronic and digitalelectronic detectors) have been used with varying degrees of success inthese fields. In general, many compromises have been made in the variousimaging and non-imaging parameters of detectors in developingoperational systems.

More recently, there has been developed a different type of detectorknown as the kinestatic charge detector (KCD). In a KCD system, there isprovided an x-ray detection volume and a signal collection volume formedin a closed chamber. In the detection volume, there is generallydisposed some type of medium which will interact with x-ray radiation toproduce secondary energy. The medium is generally enclosed within adefined space and the collection volume is preferably a multi-elementdetector of secondary energy located at one boundary of the detectionvolume. An applied electric field across the detection volume imparts aconstant drift velocity to secondary energy particles or charges drivingthe charges of one sign towards the signal collection volume. Charges ofthe other sign will drift in a direction away from the collection volumeand will not contribute to any output signal.

In the operation of the system, an x-ray beam scans a patient and thex-ray radiation passing through the patient is directed into thedetection volume. The KCD is oriented such that the one-dimensionalarray of collector electrodes spans the fan beam which is transverse tothe direction of scan, and the width of the x-ray beam in the scandirection is matched to the height of the detection volume. The x-rayradiation collides with particles in the medium of the detection volumecreating a secondary energy. The electric field across the detectionvolume is produced between a first electrode at one side of thedetection volume and the plane of the collection volume (collectionelectrodes) and the direction of the field is substantiallyperpendicular to the path of the radiation admitted into the detectionvolume. The electric field causes charge carriers between the firstelectrode and the collection electrode to drift toward the collectionelectrode at a substantially constant drift velocity. The chamberitself, including the detection and collection volumes, is mechanicallycoupled to apparatus which moves the chamber in a direction opposite tothe direction of drift of the charges at a constant velocity of amagnitude substantially equal to the magnitude of the drift velocity ofthe charges. The currents flowing in the plural collection electrodesresulting from charges produced on the collection electrodes by thecharge carriers is sensed. The spatial distribution in two dimensions ofthe radiation admitted into the chamber is determined in response to theamplitude with respect to time of the sensed current flowing in therespective plural collection electrodes. The spatial distribution ofradiation in the transverse direction is determined by the spacing ofthe collection electrodes. Thus, in a KCD system, two-dimensionalinformation can be obtained using a one-dimensional array of collectorelectrodes.

Since the motion of the chamber is at the same velocity but in adirection opposite to the drift of the charge carriers created in themedium in the detection volume, the x-ray radiation passing through eachsmall area of the patient in the x-ray beam is integrated over the timethat it takes for the charge carriers in the detection volume to driftthrough the space of the volume. This integration, which is required inorder to obtain adequate signal levels, was achieved in prior artfan-beam systems using two-dimensional arrays of collector electrodescomprised, by way of example, of 80 to 100 elements in the scandirection and 2000 elements in the transverse direction. The KCD systemprovides the same information using a one-dimensional detector array andthus avoids the cost and complexity of large two-dimensional detectorarrays as in the prior art.

Within the detection volume, a grid separates the space between thefirst electrode and the collector volume into a drift region and acollection region. The grid shields the collector electrodes from anyinduced current caused by the charges in the drift region so that onlyions in the collection region are detected by the collection electrodes.The spacing between the grid and collection electrodes is one factoreffecting the resolution of the system. The data obtained at thecollection electrodes is digitally processed to generate an image. Inthat sense, KCD is a form of digital radiography.

Digital radiography is a general term encompassing a broad spectrum ofactivities within diagnostic medical imaging. In an early form,radiographic films were digitized in an attempt to enhance and redisplayinformation of interest. The field has evolved to its current state inwhich x-ray signals are detected electronically, converted to digitalform, and processed prior to being recorded and displayed. In somecases, film is used for archival storage while in other implementationsit is totally excluded from the process.

A primary goal of digital radiography is the removal of interferingeffects from uninteresting structures in an image so that clinicallysignificant details can be displayed with enhanced visibility. Thisprocess can simplify and extend the accuracy of diagnostic procedures.Two types of subtraction techniques have been developed to accomplishthis goal: temporal or mask mode subtraction, and energy or spectralsubtraction. Temporal subtraction is used primarily in angiography,while energy subtraction has applications both in angiography andgeneral radiography.

In temporal subtraction, images are typically acquired before and afterintravenous injection of an iodinated contrast medium. These images arethen subtracted and enhanced in a digital processor to yield an image ofarteries that is devoid of shadows due to bone and surrounding softtissue. If the patient moves between the time that the mask and contrastimages are obtained, artifacts are introduced into the subtracted image,possibly interfering with the diagnostic utility of the study.

Energy subtraction is based on the fact that x-ray attenuation is anenergy-dependent phenomenon and, moreover, that the energy dependence isdifferent for materials having different average atomic numbers. Inenergy subtraction, images are acquired using different x-ray spectra,digitized, and combined in a digital processor to selectively suppresssignals due to some material or enhance signals due to others. Thistechnique can be used to image either administered or inherent contrastdifferences. In dual-energy imaging, x-ray attenuation data are obtainedusing two different x-ray beam spectra. These data can be combined in avariety of ways, each of which produces an energy subtracted image inwhich signals from a material of a specific atomic number have beeneliminated. This process is therefore referred to as material selectiveimaging. In general radiographic applications it is useful for removingunwanted objects from an image. By way of example, bone shadows can besuppressed when lung nodules are being studied in chest radiography.When the images at the different x-ray spectra are acquired closetogether in time, this method is relatively insensitive to patientmotion. There are, however, other limitations, such as residual boneshadows in tissue-cancelled images, that can interfere with thevisualization of iodinated arteries in vascular imaging applications.

In general radiography, it is desirable to eliminate shadows or imagesdue to competing anatomy. If the area of interest is lung tissue, forexample, the image shadows due to intermediate bone structures mayobscure the image of the tissue. When single energy imaging isperformed, it is extremely difficult to separate an image of a specificmaterial. However, dual energy imaging permits cancellation orseparation of image signals created by specific materials.

Before describing the present invention, a prior art system forperforming both temporal and dual energy imaging will first be describedalong with the characteristics of such a system. FIG. 1 illustrates thebasic functional components of a digital fluorography (DF) system whichis used primarily for vascular imaging applications. The imageacquisition signal chain utilizes a standard x-ray system generator,x-ray tube and image intensifier systems. The DF portion of the systemstarts at the output of the image intensifier where the image isoptically coupled to a TV camera. Both analog and digital processing ofthe video signal can be used to enhance the image prior to displayand/or image storage.

The arithmetic manipulation of images is central to the concept of DF.The ability to integrate images to reduce noise, to subtract pre- andpost-contrast images, to enhance contrast, and to reprocess must beprovided. In addition, a system controller is required to interface withthe operator as well as to keep track of the various activities withinthe system. The first alternative is to use a minicomputer for both thecontrolling and arithmetic functions. The second is to use a distributedprocessor architecture with separate arithmetic and controllingelements. The drawback of minicomputers is that they cannot handle thedata rates involved in DF if image rates greater than a few per secondare to be used. The data and computational tasks for image rates up to30 images/sec. can be handled by special-purpose hardware. Thesystemcontrolling functions are easily handled by a microprocessor.

Once the processing architecture is defined, the number of memories,their matrix size, and the number of bits per pixel need to be chosen.It is sometimes useful to integrate several frames to form a lower noiseimage even though the effective exposure time is lengthened. Memorydepth greater than the number of bits in the incoming digitized video isrequired to perform this. Although it is possible to implement temporalsubtraction using a processor with a singleframe memory, it isadvantageous to have two, full-size memories since this configurationallows frames to be integrated for both the pre- and post-contrastimages without a compromise in spatial resolution or precision.

A reasonable computational subsystem thus consists of a processor, acontroller, and a programmable special purpose arithmetic processor.

Misregistration between the mask and the image in which the arteries ofinterest are opacified is one of the principal limitations of temporalsubtraction DF.

One means of addressing the misregistration issue is to allow formationof alternate difference images. After the data are acquired, asubtraction is formed retrospectively between the contrast-filled, liveimage and an alternate mask image-one which better represents theorientation of surrounding structures in the image of interest. Thisprocedure of selecting an alternate mask after completion of theinjection procedure is called remasking.

A second means for addressing misregistration artifacts is dual energysubtraction. In digital fluorography, an X-ray image intensifier tube isused to obtain the image and it is viewed with a video camera whosesignals are digitized and stored as an image frame. After the relativelylow energy image is obtained, another image is obtained with acomparatively higher voltage applied to the X-ray tube and a resultinghigher average energy spectral band. For ordinary tissue studies the twoimages may be made in the absence of any contrast medium. Forarteriographic studies, the two images are obtained when there is anX-ray contrast medium such as an iodinated compound present in the bloodvessels.

In any case, the high average energy image picture element (pixel) dataare subtracted from the low average image data and a difference imageremains. Prior to subtraction, the data are usually variously weightedor scaled to bring about cancellation of soft tissue. The data could bescaled to reduce bone, too. However, it is not possible to remove orcancel bony structures without also removing most of the iodinatedcontrast medium which is really what one is trying to visualize inarteriographic studies.

There are also brightness non-uniformities in the subtracted ordifference image due to several effects when the data are acquired usingan image intensifier. Veiling glare, which is like haze, results fromlight diffusing or feeding back from areas of the input fluorescentscreen of the intensifier to other areas. The fact that rays of a broadX-ray beam are scattered by body tissue in an energy dependent mannerbetween ray paths also causes loss of image contrast. Differentialdetection of X-rays at various energies in the input phosphor of theimage intensifier leads to additional brightness non-uniformities. Noneof these phenomena can be completely nullified by energy subtractionalone.

A third technique for eliminating motion artifacts in a DF is the hybridsubtraction method described by W. R. Brody in U.S. Pat. No. 4,445,226with implementation alternatives described by G. S. Keyes et al. in U.S.Pat. No. 4,482,918. The hybrid subtraction method uses a combination ofenergy and temporal subtraction techniques. In hybrid subtraction, X-rayimages are obtained at two different average X-ray energies, that is,with two different kilovoltages applied to the tube and the images arecombined in a manner to suppress signals due to soft tissue in aheterogeneous object such as the body.

At this juncture it should be noted that the X-ray beams having low andhigh average energies or energy spectral bands can be obtained invarious ways. One way is by applying a constant kilovoltage (kV) to theX-ray tube and interposing two different filters alternatingly in thebeam. One filter is for softening the X-ray beam, that is, for removinghigh energy spectra above a low energy average energy band. Typically, adesired low energy spectral band is determined and a filter is chosenthat has relatively low attenuation at X-ray energies below its k-edgeand has high attenuation for energies above the k-edge to thereby removesuch high energy spectra. A filter made of a rare earth element such ascerium or erbium are examples. The other filter is for hardening thehigh energy beam and would be composed of a material that attenuates orabsorbs the low energy band intensely. Thus, the high energy spectrafilter can be aluminum, copper or brass, as examples.

Another way to generate low and high average energy X-ray beams is toswitch the X-ray tube applied voltages between low and high levels.Still another way is to switch the X-ray tube applied voltage and switchfilters correspondingly. This is the preferred way.

In hybrid subtraction a mask image is obtained first by projecting a lowaverage energy X-ray beam (hereafter called low energy beam or lowenergy spectral band) through the body followed by a higher averageenergy X-ray beam (hereafter called high energy beam or high energyspectral band) when the intravenously injected X-ray contrast medium hasnot yet entered the blood vessels in the anatomical region of interest.The images, consisting primarily of bone and soft tissue acquired at thetwo energies, are scaled or weighted, using appropriate constants, andthen subtracted to produce a mask image in which signals due to softtissue variations are suppressed and bony structures remain. The datafor a pair of high and low energy X-ray images are next obtained whenthe intravenously injected iodinated compound or other X-ray contrastmedium reaches the vessels in the region of interest. The data for thispair of images are acted upon by the same constant weighting factorsthat were used with the first pair of images and one image in this pairis subtracted from the other such that the resulting post-contrast imagecontains data representative of bone structures plus vessels containingcontrast medium. The final step in hybrid subtraction is to subtract thedual energy post-contrast image from the dual energy pre-contrast maskimage to thereby suppress or cancel the bone structures and isolate thecontrast medium containing vessels. A major advantage of the hybridsubtraction technique over temporal subtraction alone is the reducedsensitivity to soft tissue motion artifacts because the soft tissue issuppressed or cancelled in both dual energy images.

Hybrid subtraction is a good technique for eliminating anything that mayhave moved during the time between obtaining the mask image andpost-contrast image or images. However, if there is no movement duringordinary temporal subtraction, wherein the postcontrast image is simplysubtracted from the precontrast mask image, then temporal subtractionimages can be used because they generally have a better signal-to-noiseratio (SNR) than hybrid subtraction images. A higher SNR results indisplayed images that have better contrast at a given noise level.

Scattering of the X-ray beam by the body is also considered. Scatter inan image depends on X-ray beam energy, beam path length and density ofthe object being penetrated. In the hybrid subtraction technique thescattering that results from use of a broad cross section X-ray beam isof little consequence since scatter is essentially the same for eachenergy subtracted pair of images. Hence, scatter effects on imagebrightness non-uniformities are subtracted out when the pairs aresubtracted.

To recapitulate, hybrid digital fluorography techniques provide themerits of soft tissue motion insensitivity, effective bone cancellation,and elimination, to the first order, of scatter and other nonlineareffects in the X-ray image intensifier and the video camera.

The prior discussion has dealt with one type of dual-energy imagingsystem which is particularly useful for vascular imaging protocols.Limitations to this dual energy approach imposed by X-ray scatter,veiling glare, and other nonlinear energy dependent effects in theimaging chain have been pointed out. An improvement, namely hybridsubtraction, has been described which deals with the limitations invascular applications where administered contrast is being imaged.However, in general radiography alternative implementation for dualenergy subtraction must be used because inherent contrast differencesbetween relatively static anatomical components are being imaged. Themajor differences are in the apparatus used for the detection of the lowand high energy X-ray images. Once these dual energy image pairs havebeen acquired and converted into digital data, techniques and apparatusdescribed above and well-known to those schooled in the art can be usedto process, display, and archive the dual energy subtracted images.

In a general radiography system, a dual energy image can be obtained byexposing a patient to two time-spaced energy beams, one at a low averageenergy and one at a high average energy. The two beams can be obtainedby changing the anode-to-cathode voltage on an x-ray tube, by using twox-ray tubes or by inserting filters into the x-ray beam to change itsaverage energy. Another technique which has been investigated is the useof two aligned detectors separated by a beam filter. The first detectormay be a low atomic number detector while the second is a high atomicnumber detector. A discussion of dual energy imaging is given in Volume156, No. 2, pages 537-540, of Radiology journal in an article entitled"Detector for DualEnergy Digital Radiography" by Barnes et al.

As will be apparent from the foregoing discussion, dual energydifference imaging, while required in order to provide a clinicallyuseful image in many instances, has been subject to numerousimplementation problems. A further example of implementationdifficulties was experienced in attempts to develop a dual energyprojection imaging device using a computerized tomography (CT) systemwith switched energy bands. In that system, X-ray energy was switched ata frequency sufficient to generate a pair of interlaced images, i.e.,adjacent scan lines of a standard image format were obtained at a highand a low average energy. Although this arrangement was successful ineliminated motion artifacts, it gave up some resolution since only halfof the scan lines were available for each image. It might also be notedthat the scan lines were adjacent rather than overlapping and thereforenot at exactly the same image point. However, they were sufficientlyclose so as not to significantly effect imaging results.

Since the KCD system is a scanning imaging device, difficulties arise ifthe low energy and high energy images are obtained serially, i.e., theobject is scanned first with a low average energy x-ray beam and thenscanned with a high average energy x-ray beam. The time between imaginga given point in the object with the low and high average energy beamswould necessarily be at least as long as the time for a single detectorto scan the field of view of the KCD system. Such a long time spanbetween the imaging of a point with the two beams is undesirable becauseof patient motion and image registration problems.

As explained above, the KCD system accumulates charges during the timewhich the detection volume passes by a point in the target. Therefore,if an interlaced scheme, i.e., a detector forms the low energy image ofa portion of the object while the x-ray beam energy is set low and thenforms the high energy image of another portion of the object when thebeam energy is high, of dual energy imaging is used, there must be adelay between the termination of the x-ray beam pulse of one energy andthe start of the x-ray beam pulse of the other energy. If this delay isless than the time it takes for an ion to drift the length of thecollection volume, part of the signal from the detector will be due toboth the low energy and high energy x-ray beams. However, if a delay isused such that the signal from the detector is not a mixture of signalsfrom the high and the low energy pulses, more than two separatedetection volumes must be used in order to obtain a uniform patientexposure for each of the x-ray beams.

SUMMARY

In accordance with one object of the present invention and in one formthereof, dual energy difference imaging is obtained in a KCD systemusing two Kinestatic Charge Detectors located on the circumference of acircle that has its center at the focal spot of an x-ray source. The twodetectors are separated a predetermined distance and move about thecircumference of the circle at a constant velocity. The gas pressuresand electric fields of each detector are adjusted so that the ion driftvelocity in each detector is equal in magnitude but opposite indirection to the motion velocity of the detectors. Radiation from thex-ray source is directed toward the detectors and controlled wherebyradiation of a first average energy is admitted to the first detectorand radiation at a second average energy is admitted to the seconddetector. The data from one detector forms a low energy image while thedata from the other detector forms a high energy image. A dataprocessing system aligns the two images and creates a difference image.In one form the radiation energy is controlled by using filterspositioned between the x-ray source and each detector. In another form,the x-ray source is switched between high and low energy states and asynchronized slotted apparatus is used to direct low energy radiation atone detector and high energy at the other. In a third form, low energyand high energy x-ray detection volumes are aligned along a single beamfrom a constant x-ray energy source.

DESCRIPTION OF THE DRAWING

For a better understanding of the present invention, reference may behad to the following detailed description taken in conjunction with theaccompanying drawing in which:

FIG. 1 is a functional block diagram of a basic digital fluorographysystem;

FIG. 2 is a simplified illustration of a kinestatic charge detectorsystem;

FIG. 3 is one embodiment of the present invention using first and secondKinestatic Charge Detectors for obtaining a dual energy image;

FIG. 4 is another embodiment of the present invention for dual energyimaging using a KCD system;

FIG. 5 is a timing diagram illustrating high energy and low energy x-rayintensities and scan results for a system such as FIG. 4; and

FIG. 6 is a representation of a KCD chamber formed as two independentfront/back chambers separated by a radiation filter for obtaining dualenergy images.

DESCRIPTION OF THE PREFERRED EMBODIMENT

FIG. 2 is a simplified illustration of a kinestatic charge detector(KCD) system of a type with which the present invention is particularlyuseful. A detailed description of a kinestatic charge detection systemcan be had by reference to the article entitled "Kinestatic ChargeDetection" by Frank A DiBianca and Marion D. Barker, published in theMay/June, 1985 edition of Medical Physics, vol. 12, #3, pp. 339-343, andin pending patent application Ser. No. 721,727 filed Apr. 10, 1985 forDiBianca. In this system, an x-ray source 10 provides a beam of x-rayradiation 12 which is collimated by passage through a slit 14 in acollimator 16. The x-ray beam is typically 8 to 10 mm wide in the planeof FIG. 2 and 350 to 500 mm wide perpendicular to the plane of FIG. 2 atthe entrance of the detector. These two directions are referred to asthe scan direction and transverse direction, respectively. The x-rayradiation passes through a patient 18 and the attenuated radiation thenenters into an ionization chamber 20 of the kinestatic charge detectionsystem. For purposes of discussion, the KCD system may be but is notlimited to use of a gas-filled ionization chamber. The chamber 20includes an ionization space 22 preferably containing a heavy gas suchas xenon in a region between a planar anode 24 and a parallel planarcollector electrode 26. A voltage source 28 is connected between theanode 24 and the collector electrode 26 to induce an electric fieldacross the space 22 in the region between the two electrodes. A parallelplanar grid 30 is also located in the space 22 adjacent the collectorelectrode 26. The grid 30 is also provided with an electrical potentialfrom the high voltage source 28.

An x-ray photon which is absorbed in the gas within the space 22typically produces a photo electron which in turn produces a number ofelectron/ion pairs in the gas. Electrons drift rapidly to the anode 24while the ions drift much more slowly to the cathode or collectorelectrode 26. Because a relatively large voltage is present on the grid,the ions accelerate through the grid and reach the collector electrode26. The number of ions which reach the collector electrode 26 can becontrolled by adjusting the voltage of source 28 so that the electricfield between the grid and the collector electrode is sufficient toassure that a continuous field is present to direct the ions toward thecollector electrode. An imaging system 32 receives signals fromcollector electrode 26 representative of the quantity and distributionof ions reaching the electrode. The imaging system 32 uses this data toconstruct an x-ray image of the patient 18. The imaging system 32includes a data acquisition system, a computer, processing electronics,electronic data storage and image presentation equipment, all of a typeknown in the art for reproducing images from digital data.

Chamber 20 is physically moved with respect to a radiation path 34 at avelocity V_(scan) having a magnitude equal to that of the velocityv_(drift) at which the charge carriers 36 in chamber 20 are drifting.The direction in which chamber 20 is moved is opposite to the directionin which carriers 36 are drifting (and is thus perpendicular to thedirection of path 34 of the incoming x-ray beam) and has the effect ofmaking the drifting charges stationary with respect to path 34. Thecharge carriers drift with respect to the chamber 20 at a constantvelocity, and chamber 20 is synchronously moved in a manner exactlyopposite to the manner in which the charge carriers drift. Therefore,the charge carriers remain stationary with respect to path 34 for aslong as the path intersects detection volume or space 22. All x-rayphotons traveling along path 34 contribute to charges in proximity tothe path. Similar integration occurs with respect to every other pathdrawn through the patient 18 while the KCD sweeps past that path.

The apparatus and control system for moving the chamber 20 are notconsidered part of the present invention. Such apparatus may comprise amechanical structure to which the chamber 20 and associated equipmentare mounted. Servo drive systems may be provided to move the chamber 20about an arc of a circle at a predetermined velocity by means well knownin the art. Both the chamber 20 and collimator 16 are rotated such thatthe x-ray radiation scans across the patient 18 who remains stationary.

Turning now to FIG. 3, there is shown a first embodiment of the presentinvention using first and second Kinestatic Charge Detectors 38 and 40.For purposes of explanation, detector 38 will be assumed responsive tox-ray radiation at a relatively low average energy and detector 40 willreceive x-ray radiation at a relatively high average energy. Acollimator 42 forms the x-ray radiation into two fan beams 44 and 46.Collimator 42 comprises a radiation impervious sheet having a pair ofslits or slots for passing the fan beams 44 and 46 and is shown incrosssection.

Before reaching the collimator 42, the beams 44 and 46 pass throughrespective filters 48 and 50. The anode-to-cathode voltage of x-ray tubesource 10 is maintained at a constant value so that the average beamenergy remains constant. The filters 48 and 50 are of a type well knownin the art and are selected such that the average energy of the x-rayradiation in beam 44 is less than the average energy of the x-rayradiation in beam 46. A second collimator 52 is provided in the regionof the detectors 38 and 40 for preventing or at least minimizing thescatter effect, i.e., preventing x-ray radiation scattered from beam 46from reaching detector 38 and radiation from beam 44 reaching detector40.

As the apparatus effects rotation of the detectors 38 and 40, along withsynchronized rotation of collimators 42 and 52 and filters 48 and 50,the fan beams 44 and 46 scan across the patient 18. The detector 38produces data for forming the low energy image and detector 40 producesdata for forming the high energy image. The data from each detector 38and 40 are processed by the imaging system 32 to yield a dual energydifference image. The general processing techniques are well known inthe x-ray art.

For purposes of simplifying the drawings, the patient or target 18 isshown as being near the x-ray source 10 and spaced from the KCD chambers38 and 40. In an actual system, the patient 18 will be much closer tothe chambers 38 and 40 and spaced from the x-ray source 10. Thecollimators, shutters and filters will be similarly repositioned whilemaintaining their respective locations with respect to each other. Thesystem as shown in FIGS.3 and 4 allows visualization of componentplacement without distortion. More particularly, the beam width A at theentrance to KCD chambers 38 and 40 is typically between 2 and 10millimeters. If drawn to scale, the target 18 would completely cover andobscure all other portions of the system.

The widths of the fan beams 44 and 46 can be chosen to optimize thesignal-to-noise ratio (SNR) and patient x-ray exposure or dosage. Fanbeam width is adjusted by changing the width of the slots in collimator42. It should also be noted that the spacing between the detectors 38and 40 should be maintained at a minimum value so as to minimize thetime between data acquisition in each detector. As the time betweenimages increases, the likelihood of introducing motion artifacts due topatient movement also increases.

An advantage of the system of FIG. 3 is its simplicity. The x-ray sourcevoltage is maintained at a constant value thus avoiding high-voltageswitching problems and synchronizing of switching with detector motion.However, separation of the energy spectrum of each beam using filters islimiting. Although the two beams have different average energies, eachbeam includes energy of a wide range. Thus, there tends to be an overlapof energy spectra between the beams. Such energy overlap reduces theability of the difference technique to subtract out artifacts orcompeting images.

An alternate embodiment of the present invention for dual energy imagesin a KCD system is shown in FIG. 4. In this embodiment, theanode-to-cathode voltage of x-ray source 10 is varied between twodifferent values at a predetermined frequency. A pair of shutters 54 and56 are each shifted in-and-out of the respective fan beams 44 and 46 atthe frequency at which the anode-to-cathode voltage is switched. Shutter54 blocks the beam 44 when the voltage is at a high value while shutter56 blocks the beam 46 when the voltage is at a low value. The shutters54 and 56 are x-ray impervious and their mechanical construction wellknown in the art. The shutters may be electromechanically controlled bymeans well known in the art. One advantage of shutters is theindependent control of each beam exposure time. In this manner, only lowenergy radiation is admitted to detector 38 and only high energy isadmitted to detector 40. If the ratio of the amount of time that theenergy is high to the time that it is low is constant, the shutters 54and 56 could be slots in a spinning disc. However, it is preferable tobe able to vary the ratio in order to optimize the signal-to-noise ratioand patient dose. Further improvement may be realized by providing beamfilters of different materials in each beam to provide more completeseparation of the high and low energy spectra.

In order to reduce "banding" in the images, the frequency of switchingand the times for which the energy is high and low must be selected suchthat every pixel in an image is exposed for the same amount of time.This will be true if:

    F=N (V.sub.scan /W.sub.L),

N=1,2,3 . . .

and

    F=K (V.sub.scan /W.sub.H), K=1,2,3, . . .

Where W_(L) is the width of the fan beam 44 at detector 38, W_(H) is thewidth of fan beam 46 at detector 40, and F is the frequency at which thex-ray energy is switched.

A timing diagram is shown in FIG. 5 illustrating high energy and lowenergy x-ray intensities and scan results for a system in which theexposure frequency and beam widths are chosen such that K=N=1. An object58 is stationary while detectors 38 and 40 scan past. As the leadingedge 60 of detector 38 intersects the leading edge 62 of object 58, theaverage x-ray energy is set at its low value as shown by the linelabeled "Low kVp". When one-half of the width of detector 38 has passedbehind object 58, the x-ray source 10 is switched to a higher averageenergy and the beam 44 is blocked. Detector 38 will therefore not detectany radiation while the average x-ray energy is high. The line labeled"High kVp" illustrates the high energy output times. When detector 38has moved a distance equal to its width, the source 10 will switch backto its low kVp state and detector 38 will again be exposed to radiation.The position of detector 38 at the start of the second exposure is shownon line 64. The detector 38 again moves through a distance equal toone-half its width while admitting x-ray radiation.

The output signal developed by detector 38 is shown as "Low kVp Signal"and is delayed by the width of detector 38, i.e., the time that it takesdetector 38 to move a distance equal to its width. The "High kVp Signal"developed by detector 40 is obtained in the same manner as the signalfrom detector 38 but follows by a time delay equal to the transitiontime for the spacing between detectors 38 and 40 and the width of thedetectors. The relative amplitude of the signals developed by the twodetectors is indicative of the intensity of the radiation admitted.

As can be seen, by switching between high and low average energyradiation in accordance with the timing diagrams of FIG. 5, each elementof the object is exposed to both high and low average energy radiationand the radiation through each element is detected by a correspondingone of the detectors 38 and 40. Accordingly, two separate images areformed very closely in time at two different average energies. Bycontrolling the spacing between the two detectors and the scan rate (andion motion rate), the temporal spacing between images can be adjusted tominimize effects of patient motion. The frequency of switching, however,should be chosen high enough that every pixel in an image results fromions created at a large range of distances from the collector electrode26. The effects of electric field non-uniformities, scattering,recombination, space charge and diffusion of the ion cloud, all of whichare related to the distance from the collector electrode 26 at whichions are created, can be reduced if both K and N, i.e., the switchingfrequencies, are greater than two.

Another embodiment of the present invention for obtaining a dual energyimage may be obtained by mounting the detectors 38 and 40 in alignmentwith a single fan beam rather than spaced to receive two different fanbeams. A filter placed between the two detectors could remove the loweraverage energy component of the beam so that the second detector woulddetect images created by the higher average energy component. Thedetectors 38 and 40 could also utilize different atomic number x-raydetection media. A further refinement is to construct a singlefront/back split KCD chamber. As with the aligned dual chambers, thisrefined system would only require a single average energy beam. FIG. 6is a representation of a KCD chamber 20 formed as two independentchambers 20A and 20B separated by radiation filter 70. The filter 70 isselected to establish an average energy in the beam reaching detectorsection 20B at a higher value than that detected in section 20A. Thefilter 70 may have a K-edge at 69.5 keV, for example. Beam 34 entersdetector 20 as shown in FIG. 2. The two chambers may be two parts of asingle KCD chamber separated by filter 70 or may be two isolatedchambers operating with different gases (ion sources) at differentpressures. The gases and pressures could optimally be selected to passhigher energy x-rays to the second or back chamber while allowing lowerenergy x-rays to interact with the gas in the first or front chamber. Alower atomic number gas, e.g., Krypton, could be used in the firstchamber at a lower pressure than Xenon gas in the second chamber.

While the invention has been described in detail in accord with what isconsidered to be a preferred embodiment, many modifications and changesmay be effected by those skilled in the art. Accordingly, it is intendedby the appended claims to cover all such modifications and changes whichfall within the true spirit and scope of the invention.

What is claimed is:
 1. A kinestatic charge detection system comprising:asource of ionizing radiation; first and second kinestaic chargedetectors each having respectively a chamber wherein each chamberincludes a window for admitting ionizing radiation from said radiationsource, a medium in each chamber ionizble by said radiation,electrically energized electrodes within each chamber for causing one ofpositive an negative ions generated in each chamber by the radiation todrift toward a corresponding electrode at a predetermined velocity;chamber moving means coupled to each chamber for moving each chamber ina direction opposite to the direction of drift of the ions in thechamber and at a velocity substantially equal thereto, said secondchamber being mechanically coupled to follow said first chambrr throughsaid radiation; collimator means positioned between said radiation andsaid detectors for collimating the ionizing radiation into at least afirst and a second beam; means for establishing a first relatively lowaverage energy level for radiation in said first beam and a said secondrelatively high average energy level for radiation in said second beam,said establishing means comprising:(i) means for switching the averageradiation energy output of said radiation source between first andsecond valucs at a predetermined frequency; and (ii) means foralternately blocking said first and second beams for causing radiationof a first average energy level to impinge on said first chamber andradiation of a second average energy level to impinge on said secondlevel; means for directing said first beam onto said first chamber andsaid second beam onto said second chamber for forming a low energy imageand a high energy image of a patient interposed in said radiation beams;and means for forming a dual energy difference image from the low andhigh energy images.
 2. The system of claim 1 wherein said blocking meanscomprises a radiation impervious disc having at least one apertureradially positioned therein, said disc being positioned between theradiation source and the patient and beig rotated at a frquency sichthat radiation at said first average energy value passes through sadiaperture to said first chamber and radiation at said second averageenergy value passes through said aperture to said second chamber.
 3. Thesystem of claim 1 wherein said blocking means comprisedelectromechanically actuated shutters synchronized with switchng of theradiation source.
 4. The kinestatic charge detection system of claim 1and including a forst filter positioned between said collimator meansand the radiation source in alignment with said first beam, said filterestablishing an average energy in said first beam at a value differentthan an average energy in the second beam.
 5. The system of claim 4 andincluding a second filter positioned between said collimator means andthe radiation source in alignment with said second beam, sadi secondfilter establishing an average energy of said second beam different fromthe average energy of said first beam.